Vinyl polymers are used in a variety of industrial applications. For example, poly vinyl alcohol (PVA) is a highly hydrophilic polymer that is used as sizing in the textile industry, as a base gel component for the cosmetics industry, as an adherent for the paper industry and as a general adhesive. The chemical formula of PVA is (C2H4O)n and the structural formula is (—CH2CH(OH)—)n. It is widely known that PVA elicits little or no host biological response when implanted in animals. For this reason PVA is also used in a variety of biomedical applications including drug delivery, cell encapsulation, artificial tears, contact lenses, and more recently as nerve cuffs. PVA has generally not been considered for use as a load bearing biomaterial, primarily because of its low modulus and poor wear characteristics. It has been reported in the literature that hydrogel modulus and wear characteristics can often be enhanced by the formation of either chemical or physical associations. Cross-linking PVA by the addition of chemical agents such as polyaldehydes, through irradiation, or by freeze-thaw cycling, has been shown to improve the durability of PVA.
Use of PVA prepared by freeze-thawing methods has been suggested for use in biomedical applications as early as 1973. U.S. Pat. No. 3,875,302 issued to Taisei Inoue on Apr. 1, 1975 described a process of preparing gelled vinyl alcohol polymers by freezing an aqueous solution of a vinyl alcohol polymer below about −5 degrees Celsius and thereafter melting the frozen solution. The process of forming cryogels by freeze-thaw cycling was also described in a 1975 chemical engineering Ph.D. thesis by N. A. Peppas at the Massachusetts Institute of Technology (Cambridge, Mass.) and in U.S. Pat. No. 5,891,826. See also U.S. Pat. Nos. 4,472,542, 5,288,503 and 5,981,826; the entire contents of all cited references, patents and patent publications are incorporated by reference herein. Because of the slow dehydration exhibited by cryogels, they have been considered for use in contact lenses. PVA has also been considered for drug release applications, especially since the freeze-thaw process does not affect protein structure. Bioadhesive PVA gels have also been considered.
It is known that the exposure of aqueous solutions of PVA polymer to ionizing radiation can produce gels (Wang, B., et al. The Influence of Polymerconcentration on the Radiation-chemical Yield of Intermolecular Crosslinking of Poly(Vinyl Alcohol) by g-rays in Deoxygenated Aqueous Solution. Radiation Physics and Chemistry, 2000. 59: p. 91-95). Irradiation of PVA results in a chemical crosslinking of the polymer chains by the formation of covalent bonds. Hydrogels may be formed by irradiation of solid PVA polymer, PVA monomer (in bulk or in solution) or PVA polymer in solution. Irradiating a hydrophilic polymer in dry form is problematic for a variety of reasons, including the formation of unstable bonds and oxygen that cannot be fully removed. Additionally, the restricted motion of the polymer chains that bear the reactive free radicals limits the effectiveness of the cross-linking. In some hydrogels, it is possible to generate a cross-linked polymer solution by starting with pure monomer. Polymerization is performed first, followed by cross-linking, which is very convenient for many polymers. However, because of the instability of the PVA monomer, this is not a viable approach for making a PVA hydrogel. For most applications, crosslinking is conducted on polymer chains that are in solution, preferably in a deoxygenated solution.
To test the biocompatibility of PVA cryogels, Oka et al., implanted the PVA into rabbit patellar grooves and demonstrated little or no host tissue response. In further experimentation, small particles of the PVA hydrogel or UHMWPE controls, 50-300 microns in diameter, were implanted into the knee joints of wister rats. The UHMWPE induced a severe tissue response while the PVA did not induce a measurable response. The PVA was also bonded to a titanium fiber mesh which promoted bony in-growth when inset into the patellar grove of the femoral heads in rabbits. Thus, the combination titanium fiber mesh/PVA implant integrated into the joint and provided a reasonable bearing surface for joint loads.
The PVA had a low frictional coefficient when opposing articular cartilage (<0.1). Thus, it is likely that this biomaterial may be useful in hemi-arthroplasty (where wear against a hard surface is not an issue). To test the biocompatibility in this application, PVA backed with a titanium mesh was placed into the load bearing region of dog femoral condyles. The material was tolerated well and induced bony in-growth for fixation. The conclusion of Oka et al. is that this composite osteochondral device (COD), is ready for more extensive investigation as a partial articular surface replacement device.
The efficiency of a dose of gamma-radiation for crosslinking PVA in deoxygenated aqueous solution as a function of polymer concentration is shown in FIG. 1. Note that there is not only an ideal dose level, but also an optimal concentration of polymer where the efficiency of crosslinking is maximized (˜30-300 g/dm3). The peak in the crosslinking efficiency at approximately 300 g/dm3 is due to increasing degradation (random scission) of the polymer chains at higher radiation doses.
The relationship between cross-linking and degradation can be understood by considering the case of irradiated solid PVA. The irradiation of solid PVA leads to main chain degradation as a result of ketone structure formation which is not due to an oxidation step via oxygen, but through keto-enol tautomerization. In keto-enol tautomerism, a simultaneous shift of electrons and a hydrogen atom occurs. Main chain scission can then occur in the backbone bearing the keto tautomer. Keto-enol degradation is thought to dominate when the concentration of the polymer limits chain movement and free radical mobility. Thus, as the concentration passes 300 g/dm3, scission becomes more prevalent.

When ionizing radiation is applied to polymer chains in solution, reactive intermediates can be formed either by direct ionization, or indirectly by interaction with reactive intermediates (hydroxyl radicals) in the aqueous solution. In dilute solution, the indirect route dominates because of the electron fraction of the solution. Thus, for polymers in solution, the indirect route will be the primary mechanism responsible for the formation of reactive intermediates and subsequently, for the generation of crosslinks or scission. Because simple gel forming hydrophilic polymers do not have functional groups capable of efficient scavenging of free electrons, they do not participate in the formation of crosslinks extensively. The real workhorse is the hydroxyl radical in the aqueous solution. Nitrous oxide, which converts the free electrons to hydroxyl radicals, is sometimes added to polymer solutions undergoing radiation induced crosslinking to improve yield. Rosiak & Ulanski showed that the dependence of gelation dose (determined by rheology) on concentration was found to have a local minimum in the neighborhood of about 20 g/dm3 (FIG. 2, from Rosiak, J. M. & Ulanski, P., Synthesis of hydrogels by irradiation of polymers in aqueous solution, Radiation Physics and Chemistry 1999 55: 139-151). The method of crosslinking can by optimized by determining the local minimum in a corresponding gelation dose versus concentration curve for a given vinyl polymer and performing crosslinking in that range of irradiation doses.
In deoxygenated solutions, when chain break precursors are carbon-centered radicals localized at the main chain, the chain scission reactions are very slow because re-combination of radicals prevails. For non-ionic polymers like PVA, under normal irradiation conditions, chain scission yield is near zero if the concentration of polymer is low enough.
Additives can be used during the irradiation process to scavenge unwanted transient products (for example, tertbutanol scavenges OH— and nitrous oxide scavenges aqueous electrons). Other additives can help identify transient reaction products (tetranitromethane helps identify polymer radicals). Spin traps (2-methyl-2-nitrosopropane) allow EPR (or ESR) studies on short-lived species. Thiols are good H+ donors and are frequently used as polymer radical scavengers. Metal ions such as Fe(II) are also known to significantly affect the kinetics and yields of radiation-induced transformations of polyacrylic acid (PAA) (for example). Accordingly, all glassware should be carefully cleaned and even treated with complexing agents such as EDTA to remove traces of metal ions when working with polyelectrolyte gels. However, metal contamination should not cause problems when working with PVA.
Oxygen should also be considered an additive that must be carefully controlled. In oxygenated solutions, carbon centered macroradicals react with oxygen to form peroxyl radicals. The kinetics of this reaction are quite rapid (practically diffusion limited at a reaction constant of 109 dm3/mol/sec). Even in a polyanionic gel, where oxygen approach is hindered by charge effects, the reaction constant is as high as 108 dm3/mol/sec. When crosslinking with oxygen present it is important to note that neither the peroxyl nor the oxyl radicals form stable bonds upon recombination. Additionally, one of the main reaction pathways leads to chain scission (see Scheme 1 below). One method is to perform the irradiation in a sealed vessel. The oxygen present will be used up and gelation will occur. Sealed vessel irradiation has been utilized to produce hydrogel dressings. One could also irradiate in an open vessel and count on the diffusion limitation to slow the transport of oxygen from the surface. In this case, a high irradiation dose rate would be advantageous. It is also possible that a natural oxygen scavenger such as vitamin E would allow irradiation in an oxygen environment while minimizing chain scission.
Physical Properties of Irradiated Poly(Vinyl Alcohol) Hydrogels
Irradiated PVA films (60Co gamma ray source, nitrogen atmosphere, dose-rate 0.0989 kGy/min, 86 kGy total dose; 10-15 wt % 78 kDa PVA in deionized water) had a tensile strength of 19.7 MPa and a strain of 609% on breaking. Compressive modulus obtained by dynamic mechanical analysis (DMA) on 10% solutions of PVA directly irradiated by electron beam in air (100 kGy total dose) yielded a 0.5 MPa storage modulus at 1 Hz. However, the samples were quite brittle.
Cross-Linking: Freeze-Thaw Cycling
Freeze/thaw cycling of PVA polymer in solution results in the formation of physical cross-linking (i.e. weak bonding through a nonpermanent “association” of the polymer chains). PVA hydrogels formed in this manner are thermoreversible and are termed “cryogels”. In general, cryogels are solid elastomers containing over 80% water which are produced when solutions of higher molecular weight poly(vinyl alcohol) (PVA) of high degree of hydrolysis are subjected to one or more freeze-thaw cycles. Such cryogels are tough, slippery, elastomeric, resilient, insoluble in water below 50 degrees Celsius, and nontoxic.
Freeze-thaw cycling of solutions of PVA polymer results in the formation of physical associations (i.e. weak bonding through an “association” of the polymer chains). PVA hydrogels formed in this manner are termed “cryogels” and are described, for example, in U.S. Pat. Nos. 6,231,605 and 6,268,405, the entire contents of which are incorporated herein by reference. Importantly, the techniques utilized to create PVA cryogels do not require the introduction of chemical crosslinking agents or radiation. Cryogels are therefore easily produced with low impact on incorporated bioactive molecules. However, incorporated molecules are limited to those that can tolerate the freeze-thaw cycles required to make the gel. Thus the resulting material can contain bioactive components that will function separately following implantation. PVA cryogels are also highly biocompatible (as are PVA “thetagels,” discussed below). They exhibit very low toxicity (at least partially due to their low surface energy), contain few impurities and their water content can be made commensurate to that of tissue at 80 to 90 wt %.
There is still some debate over the exact mechanism that drives the gelation of PVA through a freeze-thaw cycle. However, three models have been proposed to explain the physical crosslinking that occurs during the freeze-thaw cycle: 1) direct hydrogen bonding; 2) direct crystallite formation; and 3) liquid-liquid phase separation followed by a gelation mechanism. The first two steps suggest that the gel forms through a nucleation and growth (NG) phase separation, whereas the third option pictures the process as a spinodal decomposition (SD) phase separation. Hydrogen bonding will form nodes and crystallite formation will form larger polymer crystals. However both of these mechanisms will form closely connected crosslinks, with relatively small crosslinking nodes. This observation is supported by studies on the gelation mechanism of PVA. Spinodal decomposition on the other hand causes redistribution of the polymer into polymer rich and polymer poor regions followed by a gelation process which results in more distantly spaced crosslinks. It is thought that phase separation through spinodal decomposition is likely to be responsible for the improved mechanical properties of PVA after crosslinking and occurs due to a quenching of the polymer solution. During the freezing process, the system undergoes a spinodal decomposition whereby polymer rich and poor phases appear spontaneously in the homogeneous solution. This process occurs because the phase diagram of quenched PVA (and polymers in general) at certain temperatures can have two coexisting concentration phases. The polymer rich phases are therefore highly concentrated which enhances the natural (weak) gelation of the PVA.
For cryogels, the physical characteristics depend on the molecular weight of the uncrosslinked polymer, the concentration of the aqueous solution, temperature and time of freezing and the number of freeze-thaw cycles. Thus the properties of a cryogel can be modulated. However, since the material's properties change dramatically at every freeze-thaw step, control over the properties of the finished gel is somewhat limited. The thetagels described broaden the range of functionality currently provided by PVA cryogels.
In general, the modulus of the PVA cryogel increases with the number of freeze-thaw cycles. In one experimental series, thermally cycled PVA cryogels had compressive moduli in the range of 1-18 MPa and shear moduli in the range of 0.1-0.4 MPa (see Stammen, J. A., et al., Mechanical properties of a novel PVA hydrogel in shear and unconfined compression Biomaterials, 2001 22: p. 799-806).
As cryogels are crosslinked by physical and not chemical means, there is some concern about their structural stability. The modulus of PVA in aqueous solution increases with soak time in distilled water at constant temperature. In one experiment, conducted over 40 days, the modulus increased by 50%. Theoretically, during aqueous aging, the increase in strength, with the concomitant loss of soluble PVA, is the result of an increase in the order of the supramolecular packing of the polymer chains.
It is also important to understand the effects of loss of polymer over time and how that impacts the local host biological environment. It should be noted that in this example, the cryogel was only freeze-thaw cycled once, although others have shown PVA dissolution following multiple freeze-thaw cycles. In general, there is very little information about the stability of PVA cryogel modulus under repeated load cycling (fatigue).
As might be expected, the swelling of PVA cryogels at any time point decreases with increasing number of freeze-thaw cycles, indicating a densification of the PVA gel, most likely due to a higher crosslink density. In the long term, following gelation and under static conditions, the ultimate swelling ratio decreases while the modulus increases with time. In freeze-thaw processing, temperature is used to force a phase separation of the PVA solution, thus enhancing the gelation mechanism in the PVA (it should be noted that even at room temperature a solution of PVA begins to gel weakly over time).
When PVA in aqueous solution (or in aqueous/DMSO mixtures) is heated to dissolution and then frozen and thawed repeatedly, it forms a highly elastic gel. The solgel transition forms a physically (not chemically) crosslinked polymer. Thus, the crosslinking that is achieved is thermo-reversible. There is a dependence of the cryogel characteristics on the molecular weight of the uncrosslinked polymer, the concentration of the aqueous solution, temperature and time of freezing, the heating/cooling rates and the number of freeze-thaw cycles. Thus, there is a rich parameter space from which control of the mechanical properties of the PVA cryogels may be exercised. PVA cryogels exhibit very low toxicity (at least partially due to their low surface energy), contain few impurities and their water content can be made commensurate to tissue at 80 to 90 wt % and are thus generally considered to be fairly biocompatible.
Pores can increase in size with the number of freezing-thawing cycles. It is thought that the polyvinyl polymer is rejected from the ice crystals as an impurity and is progressively “volume excluded” into increasingly polyvinyl polymer rich areas. As might be expected, the pore size increases with decreasing concentration of polyvinyl polymer.
The melting point for freeze-thaw cycled cryogels in pure aqueous solutions is about 70-80° C. The melting point of a PVA cryogel in water/dimethyl sulfoxide (DMSO) solutions increases with the number of freeze thaw cycles. For a 10-30% concentration of DMSO in water, the melting point increased with an increase in freezing time. Quantifying the complex relationship between the melting point as a function of the freezing time, the concentration of DMSO, the concentration of the PVA and the number of freeze-thaw cycles is difficult. In general, the melting point increased with PVA concentration and with the number of freeze thaw cycles. In FIG. 3, the melting point variation as a function of PVA concentration and the number of freeze thaw cycles is shown for PVA in a 1% DMSO/water solution. FIG. 3 is a graphic illustration of the dependence of melting temperature on polymer concentration, with a family of curves for different numbers of freeze-thaw cycles for cryogels in 1 vol % DMSO at −40° C., where open circles represent data from gels treated with one cycle, closed circles represent data from gels treated with three cycles, open triangles represent data from gels treated with four cycles, closed triangles represent data from gels treated with eight cycles and open squares represent data from gels treated with fourteen cycles.
Because of the increased interaction between the PVA molecules and the solvent across a range of DMSO solvent concentrations (20-30% vol), the melting point of the PVA is extremely low (near or below 10° C.). In general, the melting point increases with the number of freeze/thaw cycles and increasing PVA concentration. At very high concentrations of DMSO (90%), the cryogels have a very low melting point and were transparent. After the first freeze/thaw cycle, the melting point does not change appreciably. The melting temperature of PVA cryogels in low concentration DMSO (1-5%) is independent of freezing time. However, the melting temperature of PVA in 30% DMSO is strongly dependent on freezing time. This dependence is probably due to retarded freezing in higher concentrations of DMSO. Faster freezing reduces the effects of crystal movement on the formation of cross-links. As a consequence, the melting point of PVA frozen quickly, and then held for longer periods of time (low concentration of DMSO) is lower than for PVA that does not freeze quickly (high concentration of DMSO). At higher concentrations of PVA, the melting point dependence on freezing time in higher concentration of DMSO is not as marked. However, the melting point is already very high for these samples. The highest melting points for PVA/DMSO/Water cryogels are found in gels that do not have frozen water in them during the “freeze” (40-80% DMSO).
Effect of Thawing Rate
Gel-fraction measurements of aqueous solutions of PVA demonstrate that slower thawing leads to less leachable polymer. The data corroborates the observation of a more efficient gelation process with decreasing thaw rates. The shear modulus of the hydrogel increases approximately linearly with decreasing log of the thawing rate (FIG. 4). FIG. 4 is a graphic illustration of the dependence of the shear modulus on the log of the thawing rate for PVA hydrogels formed by a single freeze-thaw cycle of 7 g/dl solution of PVA in water (data from Yamaura, K., et al., Properties of gels obtained by freezing/thawing of poly(vinyl alcohol)/water/dimethyl sulfoxide solutions. Journal of Applied Polymer Science 1989 37:2709-2718). Low thaw rates of 0.02° C./min generate cryogels with shear moduli of 10.55 kPa for a 10 g/dL concentration of PVA. No gelling occurred in samples thawed at 10° C./min. The loss of soluble polymer in aqueous media is decreased when the initial polymer concentration is high (˜12 g/dL) provided that the thawing rate is low (˜0.02° C./min).
Modulus
In general, the modulus of the PVA cryogel increases with the number of freeze-thaw cycles. The freeze-thaw effect has been exploited to generate PVA cryogels with fairly high moduli. In an experimental series aimed at determining whether PVA cryogels could be used in load bearing applications (i.e. cartilage), thermally cycled PVA cryogels had compressive moduli in the range of 1-18 MPa (at very high strain) and shear moduli in the range of 0.1-0.4 MPa. The material used in this series of experiments is Salubria™ (available from SaluMedica, Atlanta, Ga.).
Modulus Stability.
Due to the thermoreversible nature of cryogels there has been concern in the literature about the stability of the crosslinks. It has been observed that the modulus of non-cryogel PVA in aqueous solution increases with soak time in distilled water at a constant temperature. In one experiment, conducted over 40 days, the modulus increased 1.5 times. It is possible that during aqueous aging, the increase in strength with the concomitant loss of soluble PVA is the result of an increase in the order of the supramolecular packing of the polymer chains; in other words, even at moderate temperatures, there is a weak gelation process. There are significant implications in these data for long term storage of freeze-thaw gelled PVA. It is also important to understand the effects of loss of polymer over time and how that impacts the local host biological environment.
Swelling.
As might be expected, the swelling of PVA cryogels at any time point decreases with increasing number of freeze-thaw cycles, while the storage modulus of PVA increases with the number of freeze-thaw cycles. However, following gelation and under static conditions, the ultimate swelling ratio decreases while the modulus increases with time. These observations are consistent with the theory of residual soluble polymer leaching out, proposed by Lozinsky et al. The swelling dynamics of PVA cryogels typically obey the square root law (swelling ratio vs immersion time) that is characteristic of a diffusion process.
PVA gels may also be produced through thermal cycling (not necessarily with freezing) with dehydration. Such gels are potentially suitable for use in load bearing applications, specifically, for use as an artificial articular cartilage. In such applications, an artificial cartilage can be made from PVA with a high degree of polymerization (7000), which translates to an average molecular weight of 308,000 g/mol. To generate high modulus PVA from this polymer, the polymer powder is dissolved in a mixture of water and DMSO. The solution is cooled to below room temperature to obtain a transparent gel. The gel is then dried using a vacuum dehydrator for 24 hours at room temperature and then heat treated in a silicone oil bath for 1 hour at 140° C. The PVA is placed in water until maximum hydration was achieved. The water content can be controlled by varying the annealing, or heat-treating, process. The resulting PVA hydrogel can have a water content of approximately 20%, which is low.
Examination of the material properties of this thermally cycled PVA found that the material distributes stress more homogeneously than stiff biomaterials (UHMWPE) and preserves the lubrication film gap readily in simulated articular cartilage loading. The material sustained and distributed pressure in the thin film of between 1 and 1.5 MPa. In transient load tests, the PVA withstood and distributed loads of nearly 5 MPa (FIG. 5). FIG. 5 is a graphic illustration of the time course of transient stresses transmitted through samples (mass≈27 N) of various materials dropped from a height of 10 mm, where curve 1 is polyethylene, curve 2 is subchondral bone with articular cartilage, curve 3 is subchondral bone without articular cartilage, and curve 4 is a 20% aqueous PVA hydrogel; data are from Lozinsky, V. I. and Damshkaln, L. G., Study of cryostructuration of polymer systems. XVII. Poly(vinyl alcohol) cryogels: Dynamics of cryotropic gel formation. Journal of Applied Polymer Science 2000 77:2017-2023.
Oka and colleagues further examined the wear properties of thermally cycled PVA under a variety of conditions (Oka, M, et al., Development of artificial articular cartilage, Pro. Inst. Mech. Eng. 2000 214:59-68). The wear factor found in unidirectional pin-on-disk (against alumina) experiments is comparable to that of UHMWPE. However, in reciprocating tests, the wear factor is up to 18 times larger. To improve the wear properties, PVA of even higher molecular weight and cross-linked by γ-radiation (doses over 50 kGy) were used. Such treatment reduces the wear factor considerably (to about 7 times that of UHMWPE).